Image-forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.
According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field B0 whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based. The magnetic field B0 produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse), so that the magnetization performs a precessional motion about the z-axis. The precessional motion describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the spins are deflected from the z axis to the transverse plane (flip angle 90°).
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T2 (spin-spin or transverse relaxation time). The variation of the magnetization can be detected by means of receiving RF coils which are arranged and oriented within an examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis. The decay of the transverse magnetization is accompanied, after application of, for example, a 90° pulse, by a transition of the nuclear spins (induced by local magnetic field inhomogeneities) from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (dephasing). The dephasing can be compensated by means of a refocusing pulse (for example a 180° pulse). This produces an echo signal (spin echo) in the receiving coils.
In order to realize spatial resolution in the body, constant magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field B0, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The signal data obtained via the receiving coils correspond to the spatial frequency domain and are called k-space data. The k-space data usually include multiple lines acquired with different phase encoding. Each line is digitized by collecting a number of samples. A set of k-space data is converted to an MR image by means of Fourier transformation.
Inversion recovery (IR) is generally used as a magnetization preparation technique in MR imaging. In IR imaging, the longitudinal magnetization along the main magnetic field B0 is first rotated to the negative z-direction using a 180° RF pulse. The inverted magnetization is then recovered by T1 relaxation during an inversion recovery time (TI) between the inversion and an excitation RF pulse.
A known application of IR imaging is the so-called phase sensitive inversion recovery (PSIR) method which is particularly well suited for the detection and assessment of myocardial infarction. The main challenge in PSIR image reconstruction is a phase correction process to separate the intrinsic signal phase (determining the polarity of the signal) in the complex image from phase errors (referred to in the following as background phasing) which are common in an MR image. In particular, the background phasing includes effects due to off-resonance (spatial variation of the main magnetic field). In other words, phase sensitive reconstruction is used in PSIR to remove the background phasing while preserving the polarity of the desired signal.
Several approaches have been proposed for PSIR image reconstruction including calibration of the phase errors through acquisition of another image (reference image) without IR or with IR at different TIs. However, these approaches reduce data acquisition efficiency. Further, spatial misregistration between the actual and calibration scans due to patient motion can be problematic.
An alternative approach for PSIR image reconstruction is to determine the phase errors from the IR image itself using an appropriate phase correction algorithm. One such approach is the so-called reference-less acquisition of phase sensitive inversion recovery with decisive reconstruction (RAPID) method which is able to reliably restore the polarity of the magnetization without relying on a reference image (see Jinnan Wang et al., Proc. ISMRM 2013, 2077).
Myocardial triglyceride deposition is a common source of errors and misinterpretation in PSIR-based myocardial scar detection and quantification. In the phase sensitive reconstructed images, both myocardial scar and myocardial triglyceride deposition will result in hyper-intense signals which make them hardly distinguishable. The fat signal emanating from the myocardial triglyceride deposition cannot be easily suppressed using spectrally selective pulses (e.g. according to the known SPIR/SPAIR techniques) due to the natural, fast T1 recovery at the optimal delay time (TI=200˜250 ms).